Prior to this invention, surgical lasers used on corneal tissue were calibrated by first ablating some material on the surface of a polymethyl-methacrylate (PMMA) card. The resulting concave cavity on the surface of the card created a negative-diopter lens. The power of that lens was then measured by means of a lensometer. If the reading differed from a predetermined value, typically -4 diopters, the laser intensity was adjusted by a calibration factor corresponding to the difference between the lensometer reading and the desired surgical power of the laser.
This cumbersome calibration method has two major drawbacks. In the first place, the lensometer provides an approximate reading of the curvature of the ablation. If the ablation was aspheric, wherein the power at the center of the ablation was different than the power at the periphery of it, the lensometer would only give a reading close to the power reading of the central part of the ablation. In the second place, lensometers are not very accurate and exhibit typical errors of up to 6 percent between two readings of the same ablated card. The best accuracy obtainable under this procedure is 0.25 diopters centered at -4 diopters, reflecting a 12 percent margin of error.
This prior art calibrating process for surgical laser is not only cumbersome and inaccurate, but also lengthy and ill-adapted to the environment of an operating room.
A more precise method is needed in order to evaluate the consistency of the ablation over the targeted corneal surface area. Physiologically, the human vision system cannot resolve closely spaced multiple focal points. The system translates the multi-focal data as an understandable, but not necessarily accurate image. The prior art does not provide any reliable method for verifying the regularity of the laser ablative power across the entire beam.